In a typical proton therapy system used for tumor radiation treatments for example, a proton beam is generated and output from an accelerator, e.g., a cyclotron or a synchrotron, with a certain initial energy. The initial energy determines a maximum penetration depth of the proton beam and typically is 235 MeV or higher. As the proton beam travels through a beam transportation system or a beamline, the beam energy is precisely tuned through energy selection mechanisms, e.g., an energy degrader or energy slit. The transport system includes a plurality of magnets for beam redirection (bending), focusing and steering. A rotational gantry equipped with a radiation nozzle is located at the end of the beam transport system. Eventually, the beam is delivered to a treatment station and irradiated onto a patient at an energy level prescribed for the specific treatment session based on the tumor volume, geometry, location and etc.
Due to the extremely high cost for purchasing and maintaining such a radiation system, a medical facility usually uses one accelerator for a plurality of treatment stations so the high expenditure for the accelerator facilities is distributed. Although using a multi-station single-cyclotron system is effective to distribute the cost for large medical facilities, the overall cost for such a multi-gantry system may be prohibitively high for smaller facilities that may only need one treatment station. Also, some multi-station systems do not support simultaneous treatment in multiple stations. This contribute to further disadvantage that a delay at one treatment station can cause delay at the other station. With the demand for proton beam radiation therapy increasing worldwide, smaller and less expensive proton therapy systems are highly desired to increase patient access to this treatment modality.
In proton radiation systems, including single and multi-station systems, the dipole magnets (or the bend magnets) in a gantry beamline consume significant expenditure associate with manufacture, installation, control, maintenance, and space that is limited and valuable in the medical facility.
In a gantry system, source-axis distance (SAD) refers to the distance from the iso-center to the effective source location of the proton beam. Typically, the iso-center corresponds to a center of an irradiated volume. If there are no beam focusing elements between the scanning magnets and the iso-center, the effective source location is the point where the beam changes the angle, i.e. the location of the scanning magnets. Conventionally, SAD greater than 2 m is considered necessary to achieve a parallel beam translation at the patient surface, especially in a scattering-based proton beam delivery system. A large SAD dictates a correspondingly large gantry radius, while increasing number of magnets and their complexity. Thus SAD has been the primary factor that drives the overall size and the weights of the constituents in a gantry beamline, which all contribute to the nearly prohibitive cost of manufacturing, transporting, assembling, installing, maintaining and operating such a proton therapy system.
FIG. 1 illustrates the configuration of a gantry beamline 100 in a proton radiation system with the SAD greater than 2 m in accordance with the prior art. The gantry is capable of rotating 360° around the iso-center 141. The gantry beamline 100 includes the first bend magnet 101 having an orbit bend angle of 45°, and the second bend magnet 102 having an orbit bend angle of 135°. Thus, the two bend magnets operate to bend a proton beam by 90° in total, e.g., from horizontal to vertical as illustrated.
A pencil beam scanning nozzle (not explicitly shown) may be coupled to the end of the second bend magnet 102 for delivering a dose of proton radiation to a patient. There are seven quadrupoles or focusing magnets 111-117 along the beamline 100, five of which 113-117 are disposed between the two bend magnets 101 and 102. In addition, several steering and correction magnets, e.g., 121-123, are installed between the focusing magnets 111-117. Each bend magnet 101 or 102 has an orbit bend radius of approximately 1.35 m. The second bend magnet 102 has an outer radius of approximately 1.26 m. The second bend magnet can generate a maximum magnetic field of approximately 1.8 Tesla.
The SAD 131 of this proton radiation system is approximately 2.1 m, from the center of the scanning system to the iso-center 141. Typically, a beam spot size of approximately 3-4 mm can be achieved at the iso-center. The end-to-end gantry length 132 measures approximately 9 m.
The second bend magnet 102 weighs about 22 tons, each quadrupole magnet weighs about 475 kg, and each scanning magnet weights about 1000 kg. The overall weight of such a gantry system exceeds 200 tons, including the tremendously heavy structure needed for supporting the gantry beamline.
A number of approaches have been developed or proposed for achieving a lightweight and compact gantry assembly. In one leading design on the market, the scanning system is moved upstream of the last gantry bend magnet. Such a gantry system can offer reasonably small footprints, unfortunately at considerable cost of beam precision, delivery accuracy, and many other aspects of system performance and treatment quality. In another design, reduction of gantry size is achieved by sacrificing full-range rotation of the gantry. For example the gantry can only rotate 220° instead of 360° as commonly needed for a treatment.